Magnetic resonance imaging (MRI), also called nuclear magnetic resonance imaging (NMR imaging), is a non-destructive method for the analysis of materials and is used extensively in medical imaging. It is generally non-invasive and does not involve ionizing radiation. In general terms, nuclear magnetic moments are excited at specific spin precession frequencies, which are proportional to an external main magnetic field. The radio frequency (RF) signals resulting from the precession of these spins are collected using receiver coils. By manipulating the magnetic fields, an array of signals is provided representing different regions of the volume under study. These signals are combined to produce a volumetric image of the nuclear spin density of the object.
In MRI, a body is subjected to a constant main magnetic field. Another magnetic field, in the form of electromagnetic radio frequency (RF) pulses, is applied orthogonally to the constant magnetic field. The RF pulses have a particular frequency that is chosen to affect particular nuclei (typically hydrogen) in the body. The RF pulses excite the nuclei, increasing the energy state of the nuclei. After the pulse, the nuclei relax and release RF emissions, corresponding to the RF pulses, which are measured and processed into images for display.
When hydrogen nuclei relax, the frequency that they transmit is positively correlated with the strength of the magnetic field surrounding them. For example, a magnetic field gradient along the z-axis, called the “slice select gradient,” is set up when the RF pulse is applied, and is shut off when the RF pulse is turned off. This gradient causes the hydrogen nuclei at the high end of the gradient (where the magnetic field is stronger) to precess at a high frequency (e.g., 26 MHz), and those at the low end (weaker field) to precess at a lower frequency (e.g., 24 MHz). When the RF pulse in a narrow band is applied, only those nuclei which precess at that particular frequency will be tilted, to later relax and emit a radio transmission. That is, the nuclei “resonate” to that particular frequency. For example, if the magnetic gradient caused hydrogen nuclei to precess at rates from 24 MHz at the low end of the gradient to 26 MHz at the high end, and the gradient were set up such that the high end was located at the patient's head and the low end at the patient's feet, then a 24 MHz RF pulse would excite the hydrogen nuclei in a slice near the feet, and a 26 MHz pulse would excite the hydrogen nuclei in a slice near the head. When a single “slice” along the z-axis is selected; only the protons in this slice are excited by the specific RF pulse to a higher energy level, to later relax to a lower energy level and emit a radio frequency signal.
The second dimension of the image is extracted with the help of a phase-encoding gradient. Immediately after the RF pulse ceases, all of the nuclei in the activated higher energy level slice are in phase. Left to their own devices, these vectors would relax. In MRI, however, the phase-encoding gradient (in the y-dimension) is briefly applied in order to cause the magnetic vectors of nuclei along different portions of the gradient to have a different phase advance. Typically, the sequence of pulses is repeated to collect all the data necessary to produce an image. As the sequence of pulses is repeated, the strength of the phase-encoding gradient is stepped linearly, as the number of repetitions progresses. That is, the phase-encoding gradient may be evenly incremented after each repetition. The number of repetitions of the pulse sequence is determined by the type of image desired and can be any integer, typically from 1 to 1024, although additional phase encoding steps are utilized in specialized imaging sequences. The polarity of the phase encoding gradient may also be reversed to collect additional RF signal data. For example, when the number of repetitions is 1024, for 512 of the repetitions, the phase encoding gradient will be positive. Correspondingly, for the other 512 repetitions, a negative polarity phase encoding gradient of the same magnitude is utilized.
After the RF pulse, slice select gradient, and phase-encoding gradient have been turned off, the MRI instrument sets up a third magnetic field gradient, along the x-axis, called the “frequency encoding gradient” or “read-out gradient.” This gradient causes the relaxing protons to differentially precess, so that the nuclei near the lower end of the gradient begin to precess at a faster rate, and those at the higher end precess at an even faster rate. Thus, when these nuclei relax, the fastest ones (those which were at the high end of the gradient) will emit the highest frequency RF signals. The frequency encoding gradient is applied only when the RF signals are to be measured. The second and third dimensions of the image are extracted by means of Fourier analysis.
While the z-axis was used as the slice-select axis in the above example, similarly, either the x-axis or y-axis may be set up as the slice-select axis depending upon the desired image orientation and the anatomical structure of the object of interest being scanned. For example, when a patient is laying supine in the main magnetic field, the x-axis is utilized as the slice-select axis to acquire sagittal images, and the y-axis is utilized as the slice-select axis to acquire coronal images.
Regardless of the orientation of the selected scan, mathematically, the slice select gradient, phase-encoding gradient, and read-out gradient are orthogonal. The result of the MRI scan in the true domain representation k-space is converted to image display data after a 2D or 3D Fast Fourier transform (FFT). Generally, in a transverse slice, the readout gradient is related to the kx axis and the phase-encoding gradient is related to the ky axis. In 3D MRI, an additional phase-encoding gradient is related to the kz axis to acquire data in a third dimension. When the number of phase-encoding levels is smaller than a binary number, the missing data may be filled with zeros to complete the k-space so that an FFT algorithm can be applied.
In k-space, data is arranged in an inhomogeneous distribution such that the data at the center of a k-space map contains low frequency spatial data, that is, the general spatial shape of an object being scanned. The data at the edges of the k-space map contains high frequency spatial data including the spatial edges and details of the object.
In magnetic resonance imaging, for the same set of scan parameters, a shorter scan tends to reduce the signal-to-noise ratio, while a longer scan, which would have a correspondingly larger k-space map, increases the signal-to-noise ratio as well as image quality. Ideally, a fast scan with a high signal-to-noise ratio is preferred.